Apparatus and method for measuring an optical break-through in a tissue

ABSTRACT

The invention relates to a device for measuring an optical penetration that is triggered in a tissue underneath the tissue surface by means of therapeutic laser radiation which a laser-surgical device concentrates in a treatment focus located in said tissue. The inventive device is provided with a detection beam path comprising a lens system which couples radiation emanating from the tissue underneath the tissue surface into the detection beam path. A detector device generating a detection signal which indicates the spatial dimension and/or position of the optical penetration in the tissue is arranged downstream of the detection beam path.

RELATED APPLICATION

This application is a division of application Ser. No. 10/525,424 filedSep. 7, 2005, which is a National Phase entry of PCT Application No.PCT/EP03/09345, filed Aug. 22, 2003, which claims priority to GermanyApplication No. 102 39 213.7, filed Aug. 23, 2002, and GermanApplication No. 103 23 422.5, filed May 23, 2004, each of which ishereby fully incorporated herein by reference.

FIELD OF THE INVENTION

The invention relates to a device for measuring an optical break-throughcreated in a tissue, beneath a tissue surface, by treating laserradiation which a laser surgical unit focuses in a treatment focuslocated in said tissue, said device having a detection beam pathcomprising optics. The invention further relates to a method ofmeasuring an optical break-through created in a tissue, beneath a tissuesurface, by treating laser radiation.

BACKGROUND OF THE INVENTION

Treating laser radiation is used in laser surgical methods, inparticular, for operations in optical surgery. In this connection, thetreating laser radiation is focused within the tissue, i.e. beneath thetissue surface, such that an optical break-through is created in thetissue. The treating laser radiation acts through photo disruption orphoto ablation.

In the tissue, several processes initiated by the treating laserradiation occur one after the other. If the power density of theradiation exceeds a threshold value, an optical break-through willresult, which generates a plasma bubble in the tissue. Said plasmabubble grows after creation of the optical break-through due toexpanding gases. If the optical break-through is not maintained, the gasgenerated in the plasma bubble is absorbed by the surrounding tissue,and the bubble disappears again. However, this process takes very muchlonger than the forming of the bubble itself. If a plasma is generatedat a tissue boundary, which boundary may quite well be located within atissue structure as well, tissue will be removed from said boundary.Therefore, this is then referred to as photo ablation. In connectionwith a plasma bubble which separates tissue layers that were previouslyconnected, one usually speaks of photo disruption. For the sake ofsimplicity, all such processes are summarized here by the term opticalbreak-through, i.e. said term includes not only the actual opticalbreak-through, but also the effects resulting therefrom in the tissue.

For a high accuracy of a laser surgical method, it is indispensable toguarantee high localization of the effect of the treating laser beamsand to avoid, if possible, collateral damage to adjacent tissue. It is,therefore, common in the prior art to apply the treating laser radiationin a pulsed manner, so that the threshold value for the power density ofthe treating laser radiation required to cause an optical break-throughis exceeded only during the individual pulses. In this regard, U.S. Pat.No. 5,984,916 clearly shows that the spatial area of the opticalbreak-through (in this case, of the generated interaction) stronglydepends on the pulse width. Therefore, high focusing of the laser beamin combination with very short pulses allows placing of the opticalbreak-through in a tissue with in a very punctiform manner.

The use of pulsed treating laser radiation was recently established inophthalmology, particularly for correction of visual deficiencies.Visual deficiencies of the eye often result from the fact that therefractive properties of the cornea and of the lens do not cause properfocusing on the retina. In the case of nearsightedness (also referred toas myopia), the focus is located in front of the retina when the eye isrelaxed, whereas in the case of farsightedness (also referred to ashyperopia), the focus is located behind the retina.

U.S. Pat. No. 5,984,916 mentioned above, as well as U.S. Pat. No.6,110,166, describe methods of correcting visual deficiencies by meansof suitably generating optical break-throughs, so that, ultimately, therefractive properties of the cornea are selectively influenced. Amultitude of optical break-throughs are joined such that a lens-shapedpartial volume is isolated within the cornea of the eye. The lens-shapedpartial volume which is separated from the remaining corneal tissue isthen removed from the cornea by means of a laterally opening cut. Theshape of the partial volume is selected such that, with the volumeremoved, the refractive properties of the cornea are changed so as togenerate the desired correction of visual deficiency.

In order to isolate the partial volume, it is indispensable, of course,to generate the optical break-throughs at predetermined sites. U.S. Pat.No. 5,984,916 describes corresponding sensors which sense thecross-section as well as the position and intensity of the treatinglaser beam and feed a corresponding control unit, which influences thelaser treatment beam such that an optical break-through is achieved atthe desired target point.

In contrast thereto, EP 1,232,734 A1 suggests to use a wavefront sensorto determine the position of an optical break-through which has beengenerated in an eye. In this connection, said publication states thatthe bubble size can be measured by the wavefront sensor using“relatively well-known wavefront techniques”. Unfortunately, thisdocument remains silent as to how such measurement could be effected. Itdoes not convey any technical teaching, but merely states a problem tobe solved. However, since a wavefront sensor is known to be capable ofdetermining a distortion of a wavefront, it would be conceivable toachieve the object mentioned in EP 1,232,734 A1 as the problem to besolved by using a wavefront sensor to detect a deformation of thecorneal front surface created by a bubble generated inside the cornea.However, in such a conceivable approach it would have to be expectedthat the precision of measurement strongly decreases as the size of thebubble decreases, because a smaller bubble will certainly also lead to astrongly reduced deformation of the front surface of the cornea.Moreover, it would have to be expected that, for a bubble located at agreater depth, the bubble diameter indicated by this type of measurementwould be smaller than for a bubble of the same size located higher up.Thus, in such an approach an error of measurement as a function of thedepth position of the bubble would have to be expected.

The precision with which treating laser radiation can be applied to apatient's eye to correct a visual deficiency, for example, naturally hasa direct effect on the quality of the result, i.e. in the examplementioned, on the quality of an optical correction. Therefore, it is anobject of the invention to provide a device for measuring an opticalbreak-through by which an increased selectivity of the effect oftreating laser radiation can be achieved.

SUMMARY OF THE INVENTION

This object is achieved by a device for measuring an opticalbreak-through created in a tissue, beneath a tissue surface, by treatinglaser radiation which a laser surgical unit focuses in a treatmentfocus, said focus being located in the tissue, wherein said device has adetection beam path comprising optics, in that the optics coupleradiation emitted by the tissue, from beneath the tissue surface, intothe detection beam path, and in that a detector unit is arrangedfollowing the detection beam path, said detector unit generating adetection signal which indicates the spatial extent and/or position ofthe optical break-through in the tissue. The object is further achievedby a method of measuring an optical break-through created in a tissue,beneath a tissue surface, by treating laser radiation, wherein radiationemitted by the tissue, from beneath the tissue surface, is detected anda measure of the spatial extent and/or position of the opticalbreak-through is determined therefrom.

Thus, in the concept according to the invention, the spatial extentand/or position of the optical break-through is detected and, for thispurpose, radiation is analyzed which is emitted by the tissue itself,i.e. from beneath the tissue surface. The term “emitted” summarizesback-scattered, reflected or induced radiation in contrast totransmitted radiation. The method according to the invention as well asthe device according to the invention are suitable, in particular, formeasurements during treatment of transparent or semi-transparent tissue,respectively, because, in this case, up to three-dimensional structuralinformation can be obtained. Thus, the invention is suitable, inparticular, for monitoring microsurgical operations on the eye.

The detection signal can be used to control the treatment radiation innumerous ways. Thus, on the one hand, a detection signal indicating theextent of an optical break-through can be used to control laserparameters, e.g. the beam cross-section, the radiation intensity and/orthe pulse width. In this connection, manual setting by a viewer to whomthe detection signal is presented is possible as well as a partiallyautomatic or fully automatic closed-loop control. In this case,influencing the treating laser radiation may influence parameters whichform the beam as well as deflection parameters which spatially controlthe beam in the tissue. Sensing the spatial extent of the opticalbreak-through is primarily a question of influencing the parameterswhich control the beam properties, whereas a detection signal indicatingthe position of the optical break-through in the tissue is particularlysuitable for control of a deflecting unit and, thus, allows the laserbeam to be guided by means of closed-loop control.

The detection signal generated by the detector unit can preferably beused directly for closed-loop control of the treating laser radiation.In doing so, the loop is closed, by using the actual opticalbreak-through which is formed or which acts in the tissue. This is moredirect than the use of indirect measures which are obtained prior to theinteraction with the tissue. Detection of the radiation coming from thetissue enables much more exact closed-loop control of the treating laserradiation, because the optical break-through is directly detected. Thesensing of treating laser radiation parameters before entering thetissue, as hitherto suggested in the prior art, is much more indirect,because the interaction with the tissue is left unconsidered, whereasthe approach according to the invention provides direct closed-loopcontrol.

Particularly preferred for the treating laser radiation are ultra-shortlaser pulses having pulse widths of between 1 fs and 100 ps in theinfra-red or visible spectral range (400 to 1900 nm). This opticalradiation can penetrate tissue, in particular the transparent parts ofthe eye (cornea, lens, vitreous body). At high power densities, whichare only achieved in the focal point, the laser pulses triggernon-linear optical processes in the form of optical break-throughs, suchas created, for example, by multiphoton excitation or frequencyconversion. The required power density may be specific to the tissue.

The invention is suitable, in particular, for the aforementionedsurgical method provided for correction of visual deficiencies. Inaddition, the invention may also be used for further ophthalmologicaloperations or other surgical operations. These include cuts forrefractive optical surgery or for removal of enclosed foreign bodies,cuts in the cornea, cuts in the vitreous body of the eye, in the lens orin the sclera. Localized, laser-induced changes of the tissue withoutcutting, which reduce turbidity or hardening of the cornea, are also afield in which to employ the invention. Moreover, other tissues, such asthe dermis, for example, are also transparent to infrared radiation. Asuitable selection of the wavelength of the treating laser radiationwill enable diagnosis and treatment in these tissues, too. The inventionis suitable for processing of tissue in vitro as well. Thus,histological examinations can be carried out, for example, in anextracted piece of tissue. The treating laser beam may be used to carryout a histological section in a tissue region which may have beenmeasured before. The cutting properties of the ultra-short laser pulseshave an advantageous effect on the quality of the section; collateraldamage to tissue is largely excluded.

Optical break-throughs always have an effect on tissue. In most cases,an optical scattering center, e.g. in the form of a plasma bubble, isformed for at least a short time. It is therefore preferred to determinethe spatial extent and/or position of scattering centers generated bythe optical break-through.

In a particularly advantageous embodiment of the invention, theradiation emitted by the tissue may be back-scattered radiation comingfrom an illumination radiation source. Therefore, a further embodimentcomprising an illumination radiation source, which couples illuminationradiation into the tissue, is preferred. In the method according to theinvention, observation radiation is advantageously irradiated into thetissue, and radiation emitted by the tissue in the form ofback-reflection is evaluated. Illumination radiation and treatmentradiation may be derived from the same source of radiation, e.g. byusing an energy reducer which can be switched on.

As an alternative to illumination which is externally coupled in,radiation generated by the break-through itself or back-scatteredcomponents of the treating laser radiation may be used for detection.

Detection of the position of the optical break-through in a mannersuitable to guide the laser beam in the aforementioned surgical methodrequires precision which is given by a desired correction of defectivevision. Depending on the optical application, a difference in thicknessof about 12 μm between the center and the edge of a volume to beisolated can cause a change in the refractive power of the cornea of onediopter. A correspondingly exact spatial measurement of the position ofthe optical break-through is, therefore, very advantageous, inparticular, for applications in ophthalmological surgery.

In a first embodiment of the invention, an interferometric detection ofthe radiation emitted by the tissue is effected, and the devicecomprises an illumination radiation source which, together with thedetection beam path, is part of an interferometer structure.Conveniently, back-scattered radiation is detected and the informationconcerning the optical break-through generated by the treating laserradiation, in particular the localization of the optical break-throughwith respect to adjacent layers of tissue, is obtained from radiationscattered back from the location of the focus of the illuminationradiation. The amount of back-scattered radiation bears information ondiscontinuities in refractive index, which appear both at the boundaryof different types of tissue (e.g. between the stroma and the Bowman'smembrane of the cornea) as well as at the location of the opticalbreak-through both below and above the threshold to photo disruption andphoto ablation.

In a particularly convenient embodiment, such interferometer structuremay be provided, for example, as an optical coherence tomograph, whichspatially filters, in a suitable form, radiation emitted by the tissue,e.g. back-reflected or scattered illumination radiation. In this case,spatial filtering is effected such that the optical break-through islocalized with sufficient precision. If the illumination radiation isincident along an optical axis, the spatial filtering can be achievedlaterally, i.e. perpendicular to the optical axis, by means of suitablefocusing of the illumination radiation. The focus position of theillumination radiation then laterally defines the measurement volume inwhich an optical break-through is detected or measured, respectively. Inan optical coherence tomograph, the localization of the measurementvolume can be achieved in an axial direction by using short temporalcoherence for the illumination radiation. Interference then occurs onlyin the back-reflection from the measured volume, and the presence ofinterference shows that radiation was scattered back from a known depthwithin the tissue. Since the optical break-through, as alreadymentioned, is associated with the formation of a photo disruption bubbleor a plasma bubble and separates layers of tissue, the opticalbreak-through is characterized by back-scattering of illuminationradiation which is locally increased in a significant manner. Thus, theposition of the back-scattering of the radiation emitted from the areaof the optical break-through on the optical axis of detection ispreferably determined by the occurrence of interference.

In this respect, the interferometer structure preferably comprises ameasuring arm and an adjustable reference arm. The coherence length ofthe illumination radiation determines the axial resolution, becauseinterference occurs only, if the lengths of the measuring arm and of thereference arm differ by less than the coherence length of theillumination radiation. In combination with the lateral discriminationby the focusing of the illumination radiation, the effect is, in summa,a three-dimensional spatial detection of the optical break-through, withthe diameter of the focusing defining the lateral resolution and thecoherence length of the illumination radiation defining the depthresolution.

In order to scan a larger area for optical break-throughs, it isconvenient that the source of illumination radiation focus theillumination radiation in an illumination focus located in the tissue,with the position of said focus being adjustable to generate thedetection signal. In this case, the adjustment is advantageouslyeffected with respect to the position of the focus of the treating laserradiation. In order to obtain a device which is structured as simply aspossible and is as compact as possible, preferably the same opticsdirect treating the laser radiation and the illumination radiation ontothe tissue to be treated, e.g. the cornea of an eye. In most cases, thisis achieved by combining the illumination radiation and the treatinglaser radiation via an optical combiner, e.g. by using a beam splitter.The focusing optics, through which both beams then pass together,generate a focus which is located approximately at the same locationlaterally and, depending on the beam divergence before said combination,also axially.

An independent lateral displacement of the illumination radiationrelative to the treating laser radiation can be achieved by providing anadjustable deflecting unit, e.g. a scanner, upstream of said combiner.The focus position of the illumination radiation can be axially adjustedby adapting the divergence of the illumination radiation beforecombining it with the treating laser radiation. In an analogous manner,the size of the focus can be changed relative to the treating laserradiation by suitably pre-conditioning the diameter of the illuminationradiation beam.

Therefore, it is preferred, in this respect, that the illuminationradiation be coupled into a light path of the illumination laserradiation, wherein use is made of adjustable optics, by which thedivergence of the illumination radiation is changeable independently ofthe divergence of the treating laser radiation.

In an optical coherence tomograph, the observation of the opticalbreak-through is preferably effected by axial adjustment of the spatialselection of the back-scattered radiation, including the location of theoptical break-through. This allows to observe the axial extent of theoptical break-through and its position and to obtain the detectionsignal. The heterodyne detection of the back-scattered radiationeffected by the optical coherence tomograph preferably uses anillumination radiation having as large a bandwidth as possible,preferably around an average wavelength of 850 nm. The bandwidth isinversely proportional to the coherence length of the observationradiation which also determines the axial resolution of the opticalcoherence tomograph in said heterodyne detection. Using differentwavelengths for the illumination radiation and the treating laserradiation, a superposition and separation of both optical paths can beeasily effected by dichroic beam splitters.

There are, in principle, two possibilities to realize the aforementionedaxial displacement of the focus:

If the focus diameter of the illumination beam path is greater than thatof the treating laser beam, sensing can be achieved by suitably detuningthe interferometer, e.g. by adjusting the measuring arm. The axialresolution is then mainly determined by the coherence length of theillumination radiation and is typically in the order of magnitude of 10μm. The detuning range axially defines the measuring range.

b) If the focus diameter of the illumination radiation and of thetreating laser radiation are similar in size, axial sensing requiressynchronous detuning of the interferometer and of the focus position.The latter can be effected by means of the adjustable optics, by whichthe divergence of the illumination radiation can be changed withoutchanging the divergence of the treating laser radiation, beforecombining the illumination radiation and the treating laser radiation.The axial resolution of the detection of the optical break-through thendepends on both the depth of focus of the optical projection, which isinherently confocal due to the heterodyne detection, and on thecoherence length of the illumination radiation and is below theaforementioned resolution for case a).

The higher resolution achieved in case b) is, of course, beneficial alsoto lateral sensing of the area in which the optical break-through isexpected or generated, respectively. On the other hand, an increasedspeed of adjustment is required in order to sense an area of a certainsize within the same length of time.

A further possibility of spatial selection of the radiation emitted bythe tissue, from which radiation the detection signal is obtained, isthe use of confocal imaging. In a second embodiment of the invention, itis therefore preferred that the detector unit senses the radiationemitted by the tissue by means of confocal imaging, wherein the spatialextent of the optical break-through is preferably determined byadjusting the focus of confocal imaging. The device then has a structuresimilar to that of a laser-scanning microscope, which operates byreflection according to the incident-light method.

The confocal principle allows a specific volume element in the tissue tobe projected onto the detector and to suppress any radiation from thetissue which is emitted by volume elements other than said specificvolume element. In particular, there is a strong suppression ofradiation from layers deeper down or higher up. The confocalitycondition between the sensed volume element and the projection onto thedetector, which projection is usually effected through a pinhole lens,ensures that, with a suitable chromatic correction, radiation of anywavelength arrives at the detector as long as such radiation only comesfrom the selected volume element.

The confocally detected radiation may either come from the opticalbreak-through itself, being back-scattered treating laser radiation, forexample, or may be back-scattered illumination radiation. Particularlyadvantageously, the latter can be introduced into the beam path of thetreating laser beam by means of a dichroic beam splitter. In thisconnection, an axial displacement of the selected volume element isobtained by a focus adjustment in the confocal imaging, as it is knownfrom laser scanning microscopes. It is therefore preferred that thedetector unit generates the detection signal by adjusting the focus ofconfocal imaging, preferably along the ray direction of the treatinglaser radiation. For this purpose, the adjustable optics alreadymentioned for the first embodiment can be used, said optics changing thedivergence of the illumination radiation without changing the divergenceof the treating laser radiation. Using this construction, a compactstructure is achieved, because the treating laser radiation and theillumination radiation are focused in the tissue via the same optics.

Since in confocal detection the axial resolution is inseparably linkedwith the size of the focus spot, to achieve a high resolution requiresfocusing the illumination radiation in the tissue as narrowly aspossible. An axial resolution of approx. 10 μm requires a focus diameterof approx. 3 μm. This may be achieved, relatively independently of thefocusing of the treating laser radiation, by suitably illuminating theobjective which focuses the radiation in the tissue.

During confocal detection, sensing with a lateral resolution may beeffected in a known manner by suitably deflecting the illumination beamor the focus, respectively, in a lateral direction, i.e. transversely tothe optical axis.

In a third embodiment of the invention, the axial adjustment of thevolume element selected in confocal detection can be largely dispensedwith. For this purpose, the basic principle of confocal detection ismodified such that a dispersive element, e.g. optics provided with acertain dispersion, is now used for confocal imaging. If the selectedvolume element emits white light, i.e. radiation which is composed ofdifferent spectral components, such as red, green or blue light, onlyfrom a certain depth of the sample are certain spectral componentsfocused exactly into the pinhole and thus projected onto the detectorthrough the optics. For a certain volume element, components having ashorter wavelength (blue light, for example) have a focus on the opticalaxis between the pinhole and the pinhole optics, and diverge again untilthey arrive at the detector, so that only a small part of the radiationof this spectral component can pass through the pinhole and reach thedetector. Consequently, these components are very effectivelysuppressed. The same applies to components having a longer wavelength(e.g. red light), because the focus assigned to them is located behindthe pinhole, but the beams are blocked before by the pinhole. Only aspecific central spectral component (green light, for example) isprojected by the optics from the white-light emitting volume elementthrough the pinhole.

In contrast thereto, a further white-light emitting volume element,which is arranged on the optical axis between the volume element justconsidered and the optics, may guide shorter-wavelength radiationcomponents into a focus located in the pinhole. The same applies to avolume element which, as seen from the optics, is located behind thevolume element first considered. From there, only the long-wavelengthcomponents are focused exactly on the pinhole and can be detected.

According to the third embodiment, in white-light emitting orback-scattering optical break-throughs, the spectral composition of thedetected radiation consequently encodes information concerning the depthfrom which the radiation comprising the respective spectral componentcomes. The dispersive element causes a focus shift for certain spectralcomponents of the detected radiation, said shift corresponding to thedesired axial resolution. In this embodiment, the detector for receivingthe radiation is therefore preferably provided behind the pinhole withspectral selective or resolving properties, e.g. a multi-channelspectrometer. In the simplest case, the spectrometer may comprise onlytwo or three channels.

The channels of the spectrometer are then read and evaluated, and theposition, on the optical axis, of the volume element scattering oremitting the radiation can be determined from the level of theindividual signals relative to each other. Moreover, advantageously, thesize of a radiating volume element may also be determined. If a verysmall radiating volume element is present, a distinct signal will beoutput by only one channel of the multi-channel spectrometer, in atransition region, i.e. at an average volume size, two channels willshow such distinct signal, but never all three spectral channels. If,however, a very large radiating volume element is present, thespectrometer will display approximately the same signal intensity inseveral channels, for example in all three channels. The width of thespectral distribution detected by the spectrometer thus provides ameasure of the size of the volume element, with the size along theoptical axis being decisive.

A possible design of the third embodiment relates to the kind ofradiation analyzed in the detection beam path. Various alternativesexist. On the one hand, direct emission of a plasma bubble generated byan optical break-through may be used, because the plasma emits within abroad spectrum during the process of disruption. Alternatively, thealready mentioned external illumination may be resorted to, for whichpurpose, for example, a white-light LED, a thermionic emitter, asuitably broadband laser or a superluminescence diode may be used.

In an embodiment that is particularly easy to realize, severalindependently controllable sources of radiation differing from oneanother regarding their spectral properties can be used, e.g. red, greenand blue LEDs. These individual, spectrally different sources ofradiation are sequentially operated, so that the detector no longer hasto have spectral resolution. The information for the individual colorchannels is obtained in a time sequence, so that evaluation is effectedas already mentioned above. It is, therefore, preferred that the sourceof illumination radiation comprise a plurality of radiation sourceelements, which are individually operable and have different spectralproperties, so that spectrally selective sensing is obtained bysequentially operating said radiation source elements. Spectrallydifferent illumination radiation is sequentially irradiated and therecording of the radiation emitted by the tissue, in turn, occurssequentially, to achieve the corresponding spectral association.

The aforementioned embodiments, wherein the detection beam path and theillumination beam path are incident essentially on a common opticalaxis, achieve a unit having a very compact structure. Desirableinformation on structures located on the optical axis behind the opticalbreak-through, may be obtained only with great difficulty, in somecases, because the optical break-through often has a shading effect onradiation from layers of tissue located behind the break-through. Forapplications, in which detection is desired also in those areas oftissue which are located, as seen from the apparatus, behind the opticalbreak-through, it is therefore convenient for the detection beam path tohave an optical axis which is inclined relative to an optical axis ofthe treating laser radiation or the illumination radiation. Thus, theemitted radiation is detected along an optical axis, which is obliquerelative to an optical axis along which the treating laser radiation orobservation radiation is radiated into the tissue. This approach allows,for example, to determine the thickness of a cornea while, at the sametime, sensing the optical break-through. Determining the thickness inturn allows easy gauging of the dimension of the optical break-through,e.g. of a plasma bubble diameter, because the thickness of the corneaprovides a well-known measure, which is usually precisely measured, inparticular before ophthalmological operations. Thus, an automaticcalibration for measuring the optical break-through is also achieved.

The irradiation oblique to the optical axis of the treating laserradiation is suitable for all of the above-mentioned embodiments of theinvention. However, it is useful also, in particular, for a further,fourth embodiment, which further develops the principle known from slitlamps. By means of slit optics, slit illumination can be radiated, forexample, into the cornea, oblique to the optical axis of an observationbeam path. The points of intersection of the observation beam path andof the slit illumination represent a scattered-light channel from whichradiation coming from the slit illumination and scattered in theexamined tissue can pass into the observation beam path. If the opticalaxes of the slit illumination and of the observation beam path arerotated or displaced relative to one another, the point of intersectionforming the scattered-light channel will move within the examinedtissue. For example, a displacement from the back surface of the corneato the front surface of the cornea may be effected. An intensity profilerecorded in this process will then show not only optical break-throughsforming scattering centers, but also the back and front surfaces of thecornea. The known distance of these two surfaces then gauges the size ofthe detected scattered-light center, thus of the optical break-through;accordingly, a kind of “auto-calibration” is provided.

Therefore, in a fourth embodiment of the invention, the position of theoptical axis of the detection beam path is preferably adjustablerelative to the position of the optical axis of the treating laserradiation or of the illumination radiation. Adjustment of the positionbetween the optical axis of detection and irradiation allows informationto be obtained on the spatial extent of the optical break-through or ofthe interaction induced thereby, respectively.

The oblique position of the illumination beam path or of the treatinglaser beam path and of the detection beam path relative to one anothercan be employed in the most diverse modifications, in order to determinethe position and/or extent of the optical break-throughs. This makes itpossible for the optical axes of the observation beam path and of theillumination beam path to be located obliquely relative to one anotherand to be adjustable relative to one another, independently of thetreatment beam path. Alternatively, the treatment beam path can also becombined with the observation beam path via a beam splitter.

The adjustability of the position of the observation beam path and ofthe illumination beam path can be achieved in any suitable manner bymechanical means. It is particularly easy to effect if an adjustableluminous field is employed, such as that known from DE 198 12 050 A1.The adjustment of the luminous field, which is effected, for example,using a digital mirror array, allows to easily achieve the desiredadjustment without complex mechanical assemblies.

An adjustment of the optical axes can be dispensed with, if theobservation beam path senses an at least stripe- or line-shaped image ofthe scattered-light channel. While the variant comprising adjustableoptical axes required no projection onto an imaging detector, an imagesensor is required if the optical axes of the illumination beam path andof the observation beam path have fixed positions relative to eachother. This is particularly advantageous if the illumination beam pathis combined and jointly deflected with the treating laser beam in orderto generate optical break-throughs at different sites on the tissue. Theillumination radiation and the optical break-throughs are always coaxialthen. The depth information is obtained by the inclined imaging onto theimage sensor.

In a simplified variant of the fourth embodiment, an illumination beampath is dispensed with completely. Instead, the radiation generateddirectly in the optical break-through is analyzed in the detection beampath. Back-scattering of treating laser radiation also contributes tothe generation of the detection signal. Observation is effected in anoblique manner again, using an image sensor in order to obtain therequired depth information on the position of the optical break-through.

In a variant of the fourth embodiment, which enables a particularlyuniversal measurement, a scanning deflection is effected transversely ofthe optical axis of the illumination radiation. This is preferablyindependent of the illumination beam path that feeds an image sensor. Inthis case, too, the oblique observation relative to the illuminationensures that the required depth information is obtained.

Such scanning device is generally useful for the device according to theinvention and is advantageous whenever an area located around anexpected optical break-through is to be scanned. It is thereforepreferred to use a scanning device to scan the tissue.

In a fifth embodiment of the invention, a tissue can be diagnosed aswell as three-dimensionally measured and also treated, if it is tissueintended for treatment. For this purpose, there is provided a method ofmeasuring a transparent or semi-transparent tissue, wherein illuminationlaser radiation is focused in a focal point and the position of thefocal point within the tissue is changed, to which end a variabledeflection of the illumination laser radiation is effected, and whereintissue-specific signals induced by said focusing are detected andassigned to points of measurement whose location in the tissue isrespectively defined by the determined position of the focal point, andwherein points of measurement are filtered out, thus allowing todetermine the positions of boundaries and/or inclusions in the tissue.

This method can be realized in a particularly advantageous manner usinga device for measuring a transparent or semi-transparent tissue, saiddevice comprising a source of laser radiation, a deflecting unit, afocusing unit and a detector unit as well as a control unit, whichcontrols the source of laser radiation, the deflecting unit and thefocusing unit such that illumination laser radiation emitted by thesource of laser radiation is sequentially focused in a plurality offocal points within the tissue by means of the deflecting unit and thefocusing unit, said detector unit emitting tissue-specific signals,which are induced by said focusing, to the control unit, and saidcontrol unit assigns said signals to points of measurement whoselocation in the tissue is respectively defined by the position of thefocal point, and filters out points of measurement and thus determinespositions of boundaries and/or inclusions in the tissue.

Particularly advantageously, target points for a subsequent treatment ofthe tissue by means of treating laser radiation focused in the tissuecan be obtained from the filtered-out points of measurement. In thisconnection, raster-scanning of a continuous three-dimensional area inthe tissue advantageously allows to generate a complete model of thevolume and to know, in particular, the position of boundaries andinclusions. Depending on the specific individual case, it sometimes alsomay be sufficient to scan the tissue only in a two-dimensional or evenone-dimensional manner so as to obtain information required in order togenerate target points for the action of treating laser radiation. Thepositional information of the filtered points of measurement thusprovides the control unit with an optimal regime for focusing to thetarget points. The treating laser beam can then be moved within a shorttime on an advantageous line between the target points. This also allowsto effectively avoid injuries to tissue which is not to be affected;nevertheless, a site to be treated may be located very close tosensitive parts of tissue, because the points of measurement allow anexact analysis. As to the position of the points of measurement relativeto the target points, it may be advantageous, in the interest of quickmeasuring of the tissue to select the points of measurement to have adifferent step spacing than the target points, for example a greaterstep spacing. Target points are then also positioned between points ofmeasurement, because it is advantageous to focus more target points thanthere are points of measurement that have been detected.

Since the localization of the focus of the illumination laser radiationin the points of measurement is effected by suitable control of thefocusing unit (position of the point of measurement in the depth of thetissue) and of the deflecting unit (lateral position of the point ofmeasurement), each point of measurement can be assigned to anunambiguous adjustment of the focusing unit and of the deflecting unit.The spatial position of the target point is thus also defined bysuitable parameters of the focusing unit and of the deflecting unit. Inorder to achieve high precision here, a tolerance chain, which issometimes inevitable when using two separate systems, is advantageouslyexcluded by varying the treating laser radiation locally in the tissueby the same optical means by which the position of the focal point ofthe illumination laser radiation is also influenced. The thus achievableabsolute precision in placing the focus of the treating laser radiationallows applications that were previously not possible because of thedanger of inadvertently affecting closely adjacent, sensitive tissue.

Therefore, the treating laser radiation focused into the tissuepreferably also acts as a measurement beam while being formed and guidedby the same deflecting unit and the same focusing unit. The measurementand diagnosis as well as the treatment is thus conducted using the samelaser beam which, coming from the same source of laser radiation, isfocused in the tissue by the same optical means. This has the effectthat detected points of measurement with reference to the opticalbreak-throughs and desired target points are related to the samereference point as the target points. For the measurement, the power ofthe treating laser radiation is reduced, e.g. by optical means, so thatit causes laser-induced signals in its focus which allow a measurement,but do not affect the tissue. To do so, it is advantageous to provide anenergy reducer, which can be switched into the optical path of thetreating laser beam and reduces the energy of the treating laserradiation focused in the tissue such that no irreversible change occursin the tissue, but an induced signal is caused, which signal depends onthe condition or type of the tissue and is accordingly detected in thedetection beam path.

In this case, the measured points need not be identical with thosepoints at which an optical break-through is generated. However, in thisfifth embodiment, all of said points have a common reference base,because they are generated by the same optical path, coming from thesame source of radiation. A tolerance chain, such as that potentiallyoccurring when separate systems are used, is avoided.

For said measurement, the energy of the treating laser radiation, e.g.the energy of the individual laser pulses, is reduced by means of theenergy reducer to such an extent that no irreversible changes in thetissue are created in the focal point. Alternatively, use can also bemade of a property of a pulsed source of treating laser radiation toemit background radiation of strongly reduced power between theindividual laser pulses. This background radiation can be used formeasurement, and an energy reducer can be dispensed with.

The signals detected by the detection beam path are assigned to pointsof measurement which are respectively defined by the specific positionof the detection beam path, e.g. by the specific position of a focusingunit and/or deflecting unit. These signals may be stored in a memory andmay be compared with a threshold value in a subsequent comparator, saidthreshold value being fixed or being selectable as a function of theposition of the individual points of measurement. This makes it possibleto determine all those points of measurement for which treatment is tobe provided. The corresponding positional information is transmitted toa control unit which determines a corresponding course of the generationof optical break-throughs. The treating laser radiation is then movedwith its focus along a corresponding path. This allows to effectivelyavoid injuries to tissue which is not to be acted upon, even if a siteto be treated is very close to parts of tissue that have to remainuninjured. The precision achievable is in the order of magnitude of thefocus diameter and may even be below 1 μm, depending on the focusing andon the radiation wavelength. In order to generate the opticalbreak-throughs the reduction of energy is terminated.

In the device of the fifth embodiment, a measurement step with theenergy reducer switched on can be conducted on the path already traveledin order to generate the optical break-throughs, after generation of theoptical break-throughs, in which step the optical break-throughs ortheir effect, respectively, are measured. Again, a comparison with thesignals detected during the first measurement and stored in the memoryis possible here. Depending on the result of said comparison to singleor several points treating laser radiation can be applied again toarrive at a treatment which is as successful and as gentle as possible,i.e. a second treatment step is conducted. Thus, the number of steps canbe freely selected within a sequence of measurement and treatment steps.It is therefore preferred to repeatedly determine points of measurementand target points, with treating laser radiation being respectivelyapplied to the target points.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention will be explained in more detail below, by way of exampleand with reference to the Figures, wherein:

FIG. 1 shows a perspective view of a patient during a laser surgicaltreatment using a laser surgical instrument;

FIG. 2 shows a schematic representation of the laser surgical instrumentof FIG. 1;

FIG. 3 shows the focusing of a beam onto the eye of the patient in theinstrument of FIG. 2;

FIG. 4 shows a signal course during measurement of scattering centerswhich are generated during the laser surgical treatment in the eye;

FIG. 5 shows a schematic representation explaining a cutting path duringthe laser surgical treatment;

FIG. 6 shows a light source for the laser surgical instrument of FIG. 2;

FIG. 7 shows a scanning device for the laser surgical instrument of FIG.1;

FIG. 8 shows a further embodiment of a measurement device of the lasersurgical instrument of FIG. 1;

FIG. 9 shows a schematic representation of a further embodiment of thelaser surgical instrument of FIG. 1;

FIG. 10 shows a beam path in the region of the focus of treating laserradiation from the laser surgical instrument acting on the eye;

FIG. 11 shows a schematic representation of focus positions in a furtherembodiment of a measurement device for the laser surgical instrument ofFIG. 1;

FIG. 12 shows a detailed representation of the radiation directed onto adetector of the embodiment according to FIG. 11 with a pinhole;

FIG. 13 shows a functional connection between the displacement of thefocus and the wavelength in the embodiment of FIG. 11;

FIG. 14 shows a further embodiment of the laser surgical instrument ofFIG. 1;

FIG. 15 shows a signal course as obtained with the embodiment of FIG.14;

FIG. 16 shows a further embodiment similar to that of FIG. 14;

FIG. 17 shows a further embodiment similar to that of FIG. 14;

FIG. 18 shows a further embodiment similar to that of FIG. 14;

FIG. 19 shows a schematic representation of a further embodiment of thelaser surgical instrument of FIG. 1;

FIG. 20 shows an example of a measurement regime for the scanning of atissue with the laser surgical instrument of FIG. 19, and

FIGS. 21 to 22 show further examples of a measurement regime for thescanning of a tissue with the laser surgical instrument of FIG. 19.

DETAILED DESCRIPTION OF THE DRAWINGS

FIG. 1 shows a laser surgical instrument 1, which emits a treatment beam2 being directed onto the eye 6 of a patient. The laser surgicalinstrument 1 is able to generate a pulsed treatment beam 2 such that themethod described in U.S. Pat. No. 6,110,166 can be carried out. Thepulse width is in the nanosecond or femtosecond range.

A visual deficiency in the eye 6 of the patient is remedied by means ofthe laser surgical instrument 1 by removing material from the cornea ofthe eye 6 so as to change the refractive characteristics of the corneaby a desired extent. In doing so, the material is removed from thestroma of the cornea. The stroma is located beneath the epithelium andBowman's membrane and above Decemet's membrane and the endothelium.

The material removal is effected by separating layers of tissue in thestroma by focusing the high-energy pulsed treatment beam 2 of the lasersurgical instrument 1. In this case, each individual opticalbreak-through initiates a plasma bubble, so that the separation oftissue layers covers a larger area than the focus of the treatment beam2. By suitable deflection of the treatment beam 2, plasma bubbles arelined up during treatment. The lined-up plasma bubbles circumscribe apartial volume of the stroma: the material to be removed.

Due to the treatment beam 2, the laser surgical instrument 1 operates inthe manner of a surgical knife which, without affecting the surface ofthe cornea, cuts material directly inside the stroma. If the cut isguided to the surface of the cornea by generating further plasmabubbles, the material of the stroma isolated by the cutting path can belaterally pulled out of the cornea and thus removed.

The cut performed by the laser surgical instrument 1 is guided accordingto predetermined parameters, so that the removed partial volume of thestroma causes such a change in the optical properties of the cornea thata previously existing visual deficiency is corrected to the greatestpossible extent.

In order to monitor the precision of the cutting path, the lasersurgical instrument 1 is provided with a measurement device whichdetects the spatial extents and/or the positions of the plasma bubblesgenerated by the treatment beam 2 in the corneal tissue of the eye. Themeasurement allows optimum control of the treatment beam 2. Said controlmay affect both the scanning movement of the treatment beam 2 and thecontrol of the beam parameters of the treatment beam 2 with regard tothe generation of the optical break-through. The diameter of the plasmabubbles generated is of great importance in guiding the laser beamduring the scanning operation, because the distance between individualplasma bubble centers should not exceed the average plasma bubblediameter. Otherwise, there would be a gap in the line of plasma bubbles,i.e. there would still remain continuous tissue between the disruptedvolumes. Isolation of stroma material to be removed would then bepossible only with great difficulty or even not at all. In any case, theoptical result would be unsatisfactory.

However, optimizing the parameters of the treatment beam 2 also allowsto keep the plasma bubbles as small as possible. The smaller the plasmabubbles are, the finer will be the cut formed by the laser surgicalinstrument 1. This is important, in particular, if it is taken intoconsideration that usually lens-shaped partial volumes are to be removedfrom the stroma. The precision of the cutting path and the fineness ofthe cut are particularly important at the edges of such lens-shapedvolumes.

The measured values of the measurement device may either be transformedinto a display on the laser surgical instrument 1, which allows to checkthe parameters of the treatment beam 2 or the guidance of the treatmentbeam 2 as to whether desired and pre-defined values are actuallypresent. As an alternative, it is possible to effect closed-loop controlof the parameters of the treatment beam 2 or of the deflection of thetreatment beam 2 by means of an automatic process control to obtaincertain values. Such laser surgical instrument 1 works fullyautomatically.

FIG. 2 shows a first embodiment of the laser surgical instrument 1. Themeasurement device is realized here as an optical coherence tomograph 3and follows the principle described in the publication Arch. Ophtalmol.,Vol. 112, p. 1584, “Micrometer-scale Resolution Imaging of the InteriorEye in vivo with Optical Coherence Tomography” by J. Izatt et al.

The optical coherence tomograph 3 is integrated into the treatment beam2 of the laser surgical instrument 1 via a coupling-in beam splitter 4and comprises a measuring arm 3 a as well as a reference arm 3 b. Thecoupling-in beam splitter 4 couples radiation from a measuring laser 5into the light path of the laser surgical instrument 1 such that theoptical axes of the treatment beam and of the measurement laser beamcoincide. The focus of the treatment beam 2 and the focus of theradiation of the measurement laser 5 are located in approximately thesame position laterally (not visible in FIG. 2). Both foci are formed inthe cornea of the eye 6 by focusing optics 13 arranged following thecoupling-in beam splitter 4.

FIG. 3 shows how the already combined radiation of the measurement laser5 and of the treatment beam 2 is focused into the cornea of the eye 6 asan incident beam 12 by means of the focusing optics 13. If the energydensity of the treatment beam 2 (pulse length and radiation power) isabove a certain threshold value, an optical break-through appears in thefocus, creating a plasma bubble 11.

The radiation of the measurement laser 5 is scattered back at the plasmabubble 11, coupled out again at the dichroic coupling-in beam splitter 4and then passes to a measurement beam splitter 7, where it issuperimposed on radiation from the measurement laser 5, which waspreviously split off by a measurement beam splitter 7 toward a mirror 8.The position of the mirror 8 is adjustable. Thus, the measuring arm 3 aextends from the measurement beam splitter 7 through the coupling-inbeam splitter 4 to the eye 6, and the reference arm 3 b is formedbetween the measurement beam splitter 7 and the adjustable mirror 8.

The radiation returning from the reference arm 3 b and from themeasuring arm 3 a to the measurement beam splitter 7 is brought tointerference on a photo receiver 9. Since the measurement laser 5 of useis a light source which exhibits a relatively short axial coherencelength (temporal coherence length) of about 10 μm in combination with ahigh spatial coherence, an interference pattern appears on thephotodiode 9 only if the optical lengths of the measurement arm 3 a andof the reference arm 3 b differ by not more than the temporal coherencelength of the radiation from the measurement laser 5. In order toachieve a maximum in resolution using the optical coherence tomograph 3,a measurement laser 5 having a minimal temporal coherence length, whichis below 10 μm in the embodiment example, is used.

In order to prevent any influence of the treatment beam 2 of the lasersurgical instrument 1, which is emitted by a treating laser 10 and isdirected onto the eye 6, on the optical coherence tomograph, thewavelength region of the radiation from the measurement laser 5 differsfrom the wavelength region of the observation radiation 2, and thecoupling-in beam splitter 4 has suitable dichroic properties. Forexample, a measurement laser 5 is used which emits radiation atapproximately 850 nm with a bandwidth as large as possible. A largebandwidth promotes axial resolution, because the bandwidth is inverselyproportional to the temporal coherence length of light. When usingdifferent wavelengths for the radiation of the measurement laser 5 andthe observation beam 2, superposition and separation of the beams may beeffected using a dichroic coupling-in beam splitter 4. Additionally oralternatively, suitable filters may be employed in the beam path of theoptical coherence tomography.

In order to measure the location of a plasma bubble 11, the mirror 8 isdisplaced until an interference pattern appears on the photodiode 9. Ifthe signal from the photodiode 9 which is read out by a computer PCindicates interference of the radiations from the measuring arm 3 a andfrom the reference arm 3 b, the measuring arm 3 a and the reference arm3 b have equal length, and the distance from the focus of the radiationfrom the illumination laser 5 in the cornea of the eye 11 is known. Theaccuracy of measurement is the temporal coherence length of theradiation of the measurement laser 5, i.e. the coherence length in thepropagation direction.

The back-scattering of radiation of the measurement laser 5 at the eye 6provides an indication of the discontinuities in refractive index in thecornea of the eye 6, which occur both at the boundary of differenttissues, e.g. between the stroma and the Bowman's membrane, and at thesite of interaction between the treatment beam 2 and the cornea of theeye 6, both below and above a threshold at which the above-mentionedoptical break-through occurs with photo disruption and/or photoablation.

In order to allow a maximum axial area to be measured, a device forchanging the divergence of the radiation of the measurement laser 5 (notshown in FIG. 2 for simplification) is provided between the measurementbeam splitter 7 and the coupling-in beam splitter 4, in addition to along stroke for the mirror 8, so that the divergence of the radiationcoming from the measurement laser 5 can be adjusted before coupling intothe treatment beam 2. The focusing optics 13 arranged following thecoupling-in beam splitter 4 then effect focusing of the measurement beam2 and of the radiation coming from the measurement laser 5 intodifferent focal points, wherein the adjustment of the device affectingthe divergence of the radiation coming from the measurement laser 5allows to adjust the focus position of the radiation from themeasurement laser 5, independently of the focus of the treatment beam 2.Thus, the focus of the observation radiation is adjustable along theoptical axis relative to the focus of the treatment radiation 2. Thus,back-scattered radiation also becomes detectable along the optical axisof the treatment beam 2 in an area which is greater than the stroke ofthe mirror 8.

The device for adjusting the divergence of the measurement laser 5further allows to design the spot size of the focused radiation of themeasurement laser 5 to be approximately as large as the spot size of thefocused observation beam, because scanning in an axial direction, i.e.along the axis of the observation beam 2, is then achieved bysynchronous adjustment of the divergence-changing device and the mirror8. Due to the relatively narrower focusing of the radiation of themeasurement laser 5 in this variant, the coherence tomograph 3 thusprovides an axial resolution which is even better than the temporalcoherence length of the radiation from the measurement beam 5. Theprecision of measurement accordingly increases. If this is not required,the device influencing divergence can be omitted between the coupling-inbeam splitter 4 and the measurement beam splitter 7. Axial displacementof the area from which back-scattered radiation may lead to interferenceon the photo receiver, is then effected exclusively by adjusting themirror 8, and axial resolution is mainly determined by the coherencelength of the measurement laser 5 (typically in the order of magnitudeof 10 μm).

FIG. 4 shows a signal obtained with a laser surgical instrument 1 whichcomprises an optical coherence tomograph 3 of the construction shown inFIG. 2. The signal is plotted here as a function of the adjustment ofthe mirror 8, with the stroke L of the mirror 8 being indicated as avariable in FIG. 4. The signal is derived from the interferencephenomenon detected by the photo receiver and has been generated by thecomputer PC. The signal encodes the depth of modulation of saidinterference and has a high level whenever the signal of the photoreceiver 9 indicates an interference phenomenon.

As FIG. 4 shows, a first interference appears at a displacement L1. Saidinterference is caused by radiation scattered back at the epithelium ofthe cornea. With a further displacement of the mirror 8, a second peakof the signal appears at a displacement L2. This interference is causedby the discontinuity in the refractive index at the plasma bubble 11.Said interference lasts until a displacement L3 occurs. Although thesize of the plasma bubble affects the distances L2, L3, the measurementresolution of the coherence tomograph 3 does not allow measurement ofbubble diameters of approximately 30 μm or less. The use of a source ofradiation having a shorter temporal coherence length enables suchdiameter measurement as well.

The distance between the displacement L1 assigned to the epithelium andthe beginning of the interference peak at the displacement L2 indicateshow far beneath the epithelium, i.e. the surface of the cornea, theplasma bubble is located. Thus, the position of the plasma bubble 11,caused by the treatment beam 2 in the cornea of the eye 6, can bedetected from the signal in the form of the distance from theepithelium. As an alternative to referring to the epithelium, referencemay be made, of course, to the endothelium. In this case, a new peak isto be expected, due to radiation scattered back at the endothelium, inthe representation of FIG. 4, upon further displacement. Instead of theendothelium or the epithelium, a discontinuity in the refractive indexat the Bowman's membrane or at the Decement's membrane may, of course,be used also as a reference.

The computer PC, which reads out the signal from the optical coherencetomograph 3, generates the signal. It further serves to control thetreating laser 10, for which purpose it feeds a display which indicatesthe position of the plasma bubble 11 generated by the treatment beam 2.On the other hand, this axial position of the plasma bubble is evaluatedto control the treatment beam 2 and is used to guarantee that the stepof isolating the partial volume in the stroma is effected as desired.

In an embodiment which fully automatically monitors the observance ofvalues pre-defined by treating surgeons, the computer controlsparameters of the treatment beam 2. The surgical instrument thenoperates using on-line monitoring and on-line control.

The measurement laser 5 has the structure shown in FIG. 6. It consistsof a laser 2 as well as a subsequently arranged beam-forming unitcomprising a lens 18, a stop 19 and a further lens 20. The beam diameterthus generated is adapted to the beam diameter of the treatment beam 2.By adjusting the lenses 18 and 20, the beam diameter may be additionallyreduced or enlarged in order to adjust the range of resolution ormeasurement. In addition, divergence may be changed by adjusting thelens 20, thus allowing the focus position of the measurement beam to beset.

In a schematic representation, FIG. 5 shows the cut executed by thesurgical instrument 1 under on-line control. The incident beam 12 isfocused into the cornea 14 of the eye 6 by the focusing optics 13. Asalready explained, the plasma bubble 11, which is generated by a pulseof the treatment beam 2, acts as a surgical knife which separates layersof tissue inside the cornea 14. By adjusting the incident beam 12 bymeans of a scanning unit 21, which is shown in FIG. 7 and comprisesscanner mirrors 22, 23 being rotatable about orthogonal axes, aplurality of plasma bubbles are placed next to each other, and a lenslet15 is circumscribed, which is thus cut out of the stroma of the cornea14. A possible form of cut which may achieve this is shown in U.S. Pat.No. 6,110,166, which describes a spiral-shaped line-up of plasma bubbles14.

After such cut around the lenslet 15, a lateral cut 16 is performed,which allows to remove the now isolated lenslet 15 from the cornea 14along the direction 17. The cutting path has the advantage that, in thecentral area of vision, in which the optical correction is effected byremoval of the lenslet 15, there is no cut leading through theendothelium to the outside. Instead, said lateral cut 16 is located atthe optically less important periphery of the cornea 14.

The scanning unit 21 shown in FIG. 7 serves to adjust the incident beam12. It comprises two scanning mirrors 22, 23, which are independentlyrotatable about two axes which are perpendicular to one another. Thus,the incident beam 12 can be guided across the cornea 14 in atwo-dimensional manner. Axial adjustment of the incident beam 12 iseffected by adjusting the focusing optics 13, i.e. by changing theposition of the focal point in which the energy density required for anoptical break-through, and thus the plasma bubble 11, is generated.

In order to precisely determine the axial position of the generatedplasma bubble, the mirror 8 of the optical coherence tomograph 3 isadjusted as described with reference to the signal shown in FIG. 4. Thismay be effected, for example, by an oscillation which passes through thearea between positions L1 and L2 each time. In the case of a measurementdevice having high axial resolution, such a great adjustment of themirror 8 may, however, be mechanically very complex or time-consuming insome cases. For such cases, the construction of the optical coherencetomograph 3 schematically shown in FIG. 8 is provided, which comprisesseveral, e.g. two, reference arms 3 b and 3 b′ that are substantiallysimilar to each other, i.e. they have an adjustable mirror 8 (referencearm 3 b) and 25 (reference arm 3 b′), respectively. Each reference arm 3b, 3 b′ is associated with a photo receiver 9, 26.

The length of the reference arm 3 b can be set independently of thelength of the reference arm 3 b′. Thus, radiation scattered back withhigh resolution simultaneously at the plasma bubble and at theepithelium can be measured, for example, by adjusting the reference arm3 b′ to the reflection to a reference surface, i.e. the epithelium, andreference arm 3 b detecting the boundary of the plasma bubble.

Moreover, FIG. 8 shows a possible construction of the device forchanging the divergence of the radiation from the measurement laser 5,which device was already mentioned with reference to FIG. 2, but notshown therein. The coherence tomograph of FIG. 8 is provided with anadjustable zoom optics 27.

The optical coherence tomograph 3 as used in the constructions accordingto FIG. 2 or 8 provides for a spatial selection of the area from whichback-scattered radiation is recorded. Axial selection is effected by theinterferometer construction comprising the measuring arm 3 a and thereference arm 3 b (or additional pairs thereof). The focusing optics 13effect a lateral selection by focusing. The spatial selection is freelyselectable, i.e. independently of the position of the treatment beamfocus, if the optical coherence tomograph is provided with anindependent scanning unit similar to FIG. 7. Then the area around anoptical break-through or a plasma bubble 11 can be scanned at higherresolution. The zoom optics 27 allow sensing of the area around theplasma bubble 11 not only laterally, i.e. transversely of the opticalaxis of the treatment beam 2, but also axially along said optical axis,in a large measurement range.

FIG. 9 shows a second embodiment of the laser surgical instrument 1. Thebasic principle here is to detect radiation having a large andadjustable depth of focus, which radiation is scattered back by suitablespatial filtering at the eye 6 and, in particular, at a plasma bubble 11generated there, in a manner allowing measurement of at least the axialextent of a scattering structure in the cornea of the eye, for exampleof the plasma bubble 11 generated by the optical break-through. For thispurpose, a confocal microscope 28 is provided, which operates in themanner of a laser scanning microscope.

The confocal microscope 28 is provided with a laser 29 whose radiationis adapted by an optical arrangement 30 with regard to beam parameters,such as the position of necking at the focus and the beam cross-section.The illumination radiation thus obtained is passed to a scanning unit 21by means of a splitter 31, said scanning unit comprising two scanningmirrors 22 and 23 like the scanning unit of the previous constructions.The scanning mirrors 22 and 23 are arranged closely adjacent and inimmediate proximity to a pupil of the beam path. As shown in FIG. 9, themirrors have rotary axes that are perpendicular to each other, and theymay be separately controlled. For each beam deflection which depends onthe actuation of the scanning unit 21, radiation is collected byscanning optics 32 in an aperture plane 35, from where an objective 36generates a spot image which is located in an object plane situated inthe eye 6. The treatment beam 2 is also effective in the area of thisobject plane. As in the previous embodiments, the treatment beam 2 comesfrom a treating laser 10 and is combined with the beam path of theconfocal microscope via a coupling-in beam splitter 4.

Radiation scattered back in the object plane, e.g. at a plasma bubble11, passes back to the beam splitter 31 via the described optical path.The radiation of the sensed volume element and coming from the objectplane, i.e. radiation scattered back at the eye 6, is detected in adetecting arm 42 after having passed through the beam splitter 31.Depending on the design of the beam splitter 31, the radiation isguided, at least partially, or, in the case of a dichroic design of thebeam splitter 31, almost completely, to an interference filter 38 and toa lens 39, by which a spot image of the back-scattering object in theobject plane is generated in a pinhole plane in which a pinhole aperture40 is located. The pinhole aperture 40 is followed by a photo multiplier41 which provides a signal that is transformed into an image signal byan evaluating unit (not shown) connected thereto, which unit considersthe current position of the scanning unit 21.

The size of the pinhole aperture sets the size of the object structureto be detected in the object plane. Even more decisive is, however, thata decreasing diameter of the pinhole aperture leads to an enhanced depthdiscrimination in the object plane, i.e. the size of the pinholeaperture determines from which axial area around the object planeradiation can reach the photo multiplier.

The confocal method used in the confocal microscope 28 allows to directradiation from a selected volume element in the object under examinationto the photo multiplier 41 and to almost completely suppress radiationemitted by other volume elements. In particular, radiation from tissuelayers located further down or higher up is blocked out.

The pinhole aperture 40 and the object plane from which radiation isdetected are located in conjugated planes. Adjustment of these planes,for example by adjusting the lens 39, makes it possible to effect anaxial depth scan such that the object plane from which radiation reachesthe photo multiplier 41 is adjusted. Said adjustment is required, forexample, in order to measure a plasma bubble 11 in an axial direction,i.e. along the axis on which the treatment beam 2 is incident.

Increased independence between the axial position of the aforementionedobject plane and the focal point of the treatment beam 2 may be achievedby arranging the divergence-changing device already mentioned withrespect to the constructions according to FIGS. 2 and 8, e.g. in theform of zoom optics, such that it precedes the coupling-in of thetreatment beam 2 via the coupling-in beam splitter 4. Suitableadjustment of this device then allows the object plane to be adjustedvirtually freely relative to the focal plane of the treatment beam 2.

In the construction shown in FIG. 9, a single scanning unit 21 isprovided for structural simplification, said scanning unit 21synchronously changing the lateral position of the spot image, fromwhich reflected radiation is passed to the photo multiplier 41, as wellas the focal point of the treatment beam 2. It is not possible tolaterally displace the spot within the object plane relative to thefocal point of the treatment beam 2.

Use of two independent scanning units avoids this limitation. Theillumination and detection beam path of the confocal microscope 28 canthen be independently adjusted relative to the treatment beam 2 at theeye 6. Such a construction is indicated in broken lines in FIG. 9. Thecoupling-in beam splitter 4 is then replaced by a coupling-in beamsplitter 4 a which is arranged following the scanning unit 21 of theconfocal microscope 28 in the direction of illumination. Here, thetreating laser 10 is provided with its own treatment scanner 37, whichis operable independently of the scanning unit 21 of the confocalmicroscope 28. This more complex construction allows the confocalmicroscope 28 to record radiation from the eye 6 at points which may beselected completely independently of the focus position of the treatmentbeam 2 and thus completely independently of the generation of the plasmabubbles 11. This allows not only to determine the axial extent of aplasma bubble 11, but also to measure lateral dimensions.

In the construction of the measurement device of the laser surgicalinstrument 1 shown in FIG. 9, separate observation radiation isgenerated by means of the laser 29, the optical arrangement 30 and thecoupling-in mirror 31. For some applications, this may be dispensedwith, and just a confocal detection of scattered light coming from thetreatment beam 2 at the eye 6 may be effected. In addition, aphase-sensitive detection, e.g. according to the Nomarski method, may beapplied as described in the publication by Padawer J., “The Nomarskiinterference-contrast microscope. An experimental basis for imageinterpretation.” J. Royal Microscopial Society (1967), 88, pp. 305-349,whose disclosure is explicitly incorporated by reference herein.

In the construction of FIG. 9, particularly good sensitivity isachieved, if the coupling-in beam splitter 4 or 4 a, respectively, isdichroic, i.e. if it has a beam-deflecting effect essentially only forthe wavelengths of the treatment beam 2. This dichroic property is alsoadvantageous for the variant without separate illumination radiation,because it has been shown that in plasma bubbles 11 broad-band radiationis generated also outside the spectral range of the treatment beam 2.

The confocal detection method is usually effected using optics which arechromatically corrected to the best possible extent. This ensures thatthe radiation from a specific volume element arrives at the detector,e.g. the photo multiplier 41, in a wavelength-independent manner.

In contrast thereto, according to a third embodiment modifying thesecond embodiment, the optics which project light from the objectiveonto the pinhole aperture 40 are deliberately designed to be dispersive.In this connection, FIG. 10 shows an enlarged view of the detecting arm42 of the confocal microscope 28.

An incident detection beam 43 is transformed by the lens 39, which isnow provided as a multi-component lens, into a focused detection beam44. Due to dispersive, i.e. wavelength-dependent, focusing properties ofthe optical group forming the lens 39, the effective focal length forlight in the blue spectral region is clearly shorter than for light inthe red spectral region. Thus, white light, which consists of red, greenand blue spectral components and is emitted by a volume element, isfocused into different foci, as shown by way of example in FIG. 11. Theblue radiation 48 is focused into a focal plane 45, which is located infront of a focal plane 46 for green radiation, which is in turn followedby a focal plane 47 for red radiation 50.

The dispersive properties of the lens 39 are utilized for spectrallyselecting depth information or for obtaining further depth information.As shown in FIG. 12, for a given volume element, the dispersive opticsof the lens 39 focus only green spectral components precisely in thepinhole aperture 40, because only for these spectral components is thefocal plane 46 located in the plane of the pinhole aperture 40. Thefocus of the blue radiation 48 is located on the optical axis in frontof the plane of the pinhole aperture 40, so that the blue radiationdiverges again on the way between its focal plane 45 and the plane ofthe pinhole aperture 40. Thus, only a small, almost unnoticeable part ofblue radiation 48 is transmitted through the pinhole aperture 40,resulting in effective suppression of the blue radiation 48 coming fromthe given volume element. This likewise applies for the red radiation50, because the focal plane 47 thereof is located behind the plane ofthe pinhole aperture 40; the red radiation from the volume element isthus also blocked by the pinhole aperture 40. In conclusion, in theschematic representation of FIG. 12, only green radiation 49 of thevolume element reaches the detector, which is provided as a spectrometer51 here.

Referring now to a further white-light radiating volume element, whichis located on the optical axis between the previously considered volumeelement and the optics, the green component thereof is now focused in aplane behind the plane of the pinhole aperture 40. However, the blueradiation from this further volume element has a focus in the pinholeaperture, if the axial distance of both volume elements correspondsexactly to the focus displacement, with the magnification of the opticalprojection having to be taken into consideration as well. The sameapplies to a volume element, which, as seen from the objective, isaccordingly located behind the first explained volume element. From thisthird volume element, only red components of the radiation can passthrough the pinhole aperture 40. Thus, a spectral radiation mixturepasses through the pinhole aperture 40, wherein the spectral informationencodes the axial position of the volume element emitting saidradiation.

The spectral analysis of the radiation passing through the pinholeaperture 40 by means of the spectrometer 51, therefore, provides axialresolution, without having to adjust the confocal microscope. Thespectral information carries depth information about the volume element,from which the radiation of the spectral region comes. The axialposition of the volume element, from which the radiation of the spectralchannel originates, is determined from the relative height of individualspectral channels. In the simplest case, an evaluation using two orthree channels will suffice.

Further, the size of a radiating volume element is determined in thisapproach. In a three-channel spectral analysis a small radiating volumeelement will generate, in only one channel, a signal exceeding a certainthreshold value. For a medium-sized volume, two signals may alsoindicate a corresponding signal, but never all three channels, as longas the volume element is smaller than the focus offset between thosechannels that are spectrally the furthest apart. If there is a largeradiating volume element, however, the spectrometer will essentiallyprovide approximately equal signals in all channels. The width of thespectral distribution, which is indicated by the spectrometer 51, thusbears information on the size of the volume element. The wider thespectral distribution is, the larger is the volume element.

The focus offset dF achieved by the dispersive optical group of the lens39 is plotted in FIG. 13 as a function of the wavelength LAMBDA, withall values being indicated in μm. For a spectral region from blue tored, there will be a focus displacement in the order of magnitude of atleast 1 mm. Thus, the use of the aforementioned spectrally selectivelydetecting confocal microscope 28 allows detection of plasma bubbleshaving a maximum extent of up to 1 mm. Since this range of measurementusually suffices, the third embodiment can also dispense with an axialadjustment of the object plane of the confocal microscope and stillprovide information on the axial extent of a plasma bubble 11 sufficientfor on-line control.

Of course, the spectral distribution of the radiation coming from theexamined object also plays a role here. The spectral detection ofconfocally recorded radiation when using dispersive optics requires thatradiation having a certain broad-band spectral distribution be availablefrom the object plane. In this case, in the simplest embodiment, theradiation of a plasma bubble 11 to be detected may be sensed directly,because the plasma emits, during the process of disruption which wasinitiated by the treatment beam 2, radiation having a broad spectrum.For certain applications, however, the spectral distribution of theradiation may have properties that are not sufficiently constanttemporally or spectrally. For these cases, external illumination isprovided. Said external illumination must have the required spectralbandwidth, e.g. a white-light LED or a light bulb can be used.

In a variant of the third embodiment, no spectrometer 41 is employed.Instead, illumination of the volume element to be detected or of thespot, respectively, is effected using spectrally controlled radiation.The radiation is spectrally adjusted in a sequential manner. In a simpleconstruction, a blue, a green and a red LED are used for illumination,which are sequentially switched on. Information on the individualspectral channels is then obtained in a time-sequential manner, so thata suitable design of the source of illumination in this modificationwill allow working with an inexpensive detector which exhibits no, orotherwise only insufficient spectral discrimination properties.

FIG. 14 relates to a fourth embodiment of a laser surgical instrument 1.For the measurement device, it uses the principle of a slit lamp byirradiating illumination radiation on an optical axis which extendsinclined to an optical axis along which observation takes place.

The measurement device is designed as a slit-lamp arrangement 52, whichilluminates the cornea 14 of the patient's eye 6 in a slit-shapedmanner. The slit optics 53 image a very narrow light bundle into thecornea, along an optical axis 60 of the slit illumination. Said lightbundle may be generated, for example, using a slit or a part whichserves as a controllable or programmable stop, such as a micromirrorarray or the like. In the construction of FIG. 14, a micromirror arrayis used, which is illuminated with radiation. Only one or few mirrorlines reflect the light to the cornea 14. The slit illumination via theslit optics 53 may be adjusted with regard to the angle of incidence,i.e. with regard to the position of the optical axis 60, in a scanningunit 54, which is indicated by the bent double arrow in FIG. 14. This iseffected by suitably controlling the micromirror array.

In the cornea, the input slit-shaped illumination is scattered atscattering centers, for example the epithelium, the endothelium, orplasma bubbles 11 generated by the laser surgical instrument 1. Thescattered light is recorded by an objective 56, which also focuses thetreatment beam 55 into the cornea 14. The treatment beam 55 of the lasersurgical instrument 1 is coupled into the optical path via thecoupling-in beam splitter 4 also in this embodiment. The coupling-inbeam splitter 4 separates the scattered light recorded by the objective56 and guides it to a photo receiver 59 by means of imaging optics 58.The objective 56, the coupling-in beam splitter 4 and the imaging optics58 form an observation beam path.

The photo receiver 59 only records radiation which comes from theintersection point between the optical axis 60 of the slit illuminationand the optical axis 61 of the selection or observation beam path. Ascattered-light channel 57, from which scattered light resulting fromthe slit illumination is recorded, is defined at the intersection point.

By the scanning movement of the slit illumination, the scattered-lightchannel is displaced along the optical axis 61 of the detection beampath. A rotation of the optical axis 60 of the slit illumination to theleft, in the representation of FIG. 14, displaces the scattered-lightchannel toward the endothelium, and a rotation to the right displaces ittoward the epithelium. The rotation of the optical axis 60 of theslit-shaped illumination allows scanning of the cornea 14 along theoptical axis 61, and the position of a plasma bubble 11 can be preciselydetermined by a scanning displacement of the scattered-light channel 57from the endothelium to the epithelium. The bubble's position can bereferenced by referencing to the reflections from the endothelium or theepithelium, respectively, or Bowman's membrane with regard to thedistance therefrom, thus enabling not only determination of the diameterof the plasma bubble 11, but also (absolute) position detection relativeto the endothelium and the epithelium.

The signal obtained with the slit-lamp arrangement 52 is represented inFIG. 15, which shows the signal level, i.e. the radiation intensityregistered by the photo receiver 59 in a signal course 63, plottedagainst a scanning angle α of the optical axis 60, along which theslit-shaped illumination is irradiated.

As can be seen, a first signal peak appears at an angle α0, said angleresulting from a reflection at the back-surface 64 of the cornea 14,i.e. from the reflection at the endothelium. A plasma bubble reflection65, whose width is a measure of the extent of the plasma bubble 11 alongthe optical axis 61, is detected between angles α1 and α2. Finally, atthe angle α3, a front-surface reflection coming from the endothelium isregistered. The known thickness of the cornea 14 gauges the distancebetween the angles α0 and α3 as a measure of thickness, allowing toconvert the distance between α0 and α1, i.e. the height of the plasmabubble above the endothelium, the distance between α1 and α2, i.e. thethickness of the plasma bubble, as well as the distance between α2 andα3, i.e. the depth of the plasma bubble beneath the epithelium, to ameasure of length.

If lateral information is to be obtained, in addition to the axialinformation provided by the scanning movement of the slit-shapedillumination, imaging onto an imaging detector instead of the photoreceiver 59 may be effected in the detection beam path.

This principle is shown in FIG. 16, which relates to a variant of thefourth embodiment. The construction of FIG. 16 corresponds largely tothat explained with reference to FIG. 14. Identical parts have the samereference numerals and are, therefore, not described again.

The slit-shaped illumination is now incident centrally, i.e. coaxiallyto the treatment beam 55. This provides an illumination beam path 70,whose optical axis 60 coincides with the optical axis 61 of thetreatment beam path. Observation is effected along an optical axis 61which is located obliquely to the optical axis 60 of the illuminationbeam path. Thus, the observation beam path 67 is located at an angle tothe illumination beam path, as was the case also in the embodiment ofFIG. 14. However, in the construction of FIG. 16, the scanning movementof the illumination beam path is caused by a scanning unit 54 that isprovided for the treatment beam anyway. The radiation coming from thescanning unit 54 is deviated again, in the construction shown in FIG.16, through a mirror 68 which may be optionally provided as a beamsplitter and enables observation of the field of operation through amicroscope.

The scanning unit 54 changes the angle of incidence of the optical axis60 relative to the eye 6, at which angle the illumination radiation ofthe illumination beam path 70 is incident on the cornea 14. Theobservation beam path 67 as well as its optical axis 61 are notadjustable in the embodiment of FIG. 16, although this may be optionallypossible, of course, in order to obtain additional information. Thescattered-light channel 57 is thus displaced along the (fixed) opticalaxis 61 of the observation beam path. In the observation beam path 67,an image receiver 69 is arranged following the imaging optics 58, saidimage receiver 69 recording an image of the scattered-light channel 57located at the point of intersection of the axes 61 and 60. Evaluationof the image from the image receiver 69 yields information on the sizeand the position of the plasma bubble 11, and an intensity evaluation ofthe image in a cut leading through the plasma bubble 11 or acorresponding projection provides a signal similar to FIG. 15.

FIG. 17 shows a further variant of the fourth embodiment of FIG. 16.Parts already shown in FIG. 16 are identified by the same referencenumerals and are not explained again. Like the construction of FIG. 16,the slit-lamp arrangement 52 of FIG. 17 also provides an observationbeam path 67, whose optical axis 61 extends obliquely to the opticalaxis 62 of the treatment beam path. In the observation beam path 67, theimaging optics 58 are arranged preceding an image receiver 69. However,no additional illumination radiation is coupled in now, but instead,treatment radiation scattered directly at a plasma bubble 11, orradiation generated in the plasma bubble 11 itself, is detected.Observation is effected from an inclined view, so as to derive depthinformation from the signal of the image receiver 69. In addition,visual observation by means of a microscope is possible via the beamsplitter 4, which no longer serves for coupling in now.

FIG. 18 shows a further variant of the fourth embodiment. The slit-lamparrangement 52 which serves as the measurement device in the lasersurgical instrument 1, corresponds essentially to the construction ofFIG. 17. Now an illumination beam path 70, which irradiates aslit-shaped illumination via the objective 56, along the optical axis 62of the treatment radiation 55, is coupled in via the coupling-in beamsplitter 4. The coupling-in beam splitter 4 combines the optical axis 60of the illumination radiation with the optical axis 62 of the treatmentradiation 55.

An illumination scanning unit 71 is arranged preceding the coupling-inbeam splitter 4, said unit carrying out a deflection of the slit-shapedillumination independent of the treatment radiation 55. This allows thescattered-light channel 57, from which scattered light of theslit-shaped illumination reaches the observation beam path 67, to beadjusted also laterally relative to the focus of the treatmentradiation. Thus, in addition to the information obtained by theconstruction of FIG. 16, the signal of the image receiver 69 allows toobtain information on the scattered-light image and, thus, on thestructure of the cornea laterally of the focal point of the treatmentradiation 55.

Of course, the information obtained through the measurement device ofthe third or fourth embodiment with regard to position or size of theplasma bubble 11 is used to control the laser surgical instrument, sothat open-loop and/or closed-loop on-line control are achieved here aswell.

FIG. 19 schematically shows a fifth embodiment of a laser surgicalinstrument, comprising a pulsed source of laser radiation 71 with apulse energy that is sufficiently large for treatment of the tissue inquestion; a deflecting unit 72, which laterally deflects the laser beamcoming from the source of laser radiation 71; a tunable focusing unit73, by which the position of the focal point in the depth of the tissueis set; a positioning unit 74 for positioning the tissue 6 to betreated, as well as a control unit 75 for control of the source of laserradiation 71, the deflecting unit 72 and the focusing unit 73. Thecontrol unit 75 controls the aforementioned components such that thefocus of the laser radiation is sequentially focused on real targetpoints ZP′ having the coordinates (x″, y″, z″). To this end, informationon the coordinates (x″, y″, z″) of desired target points ZP has to beavailable to the control unit 75. Depending on the application, thepositioning unit 75 may be omitted, or may be replaced by a deviceensuring coarse positioning of the tissue 76.

In the laser surgical instrument of FIG. 19, there are further providedan energy reducer 77, a detector 78 and a memory unit 79. The radiationenergy of the pulsed source of laser radiation 71 is attenuated by theenergy reducer switched into the beam path to such an extent that thefocus of the laser radiation in the tissue 76 will not cause anyirreversible changes. The laser beam 80 emitted by the source of laserradiation 71 can thus be scanned in a focused manner in the tissue 76 asa measurement beam (in a so-called regime of measurement), on the onehand, and also as a treatment beam (in a so-called regime of treatment),on the other hand. As measurement beam the laser beam 80 causes alaser-induced signal S in the real measurement point MP as a function ofthe properties of the tissue said signal S being received by a detector78 via the detection beam path (which is not shown in further detail).The detected laser radiation-induced signals S are supplied from theoutput of the detector 78 to the input of a memory unit 79 and arestored in the memory unit 79 together with the coordinates (x′, y′, z′)of the detected associated points of measurement MP′. The laserradiation-induced signals S are compared in a comparator unit 81, whichis connected to the output of the memory unit 79, with threshold valuesS stored therein.

This selects those points of measurement onto which the treating laserbeam is to be directed as target points when the energy reducer 7 isinoperative. This defines the regime of treatment. The coordinates ofthese selected detected points of measurement MP′ are transmitted to thecontrol unit 75 and are available to control the deflecting unit 72 andthe focusing unit 73.

Transformation of the coordinates of the detected points of measurementMP′ into coordinates of the desired target points is not required,because they are based on the same coordinate system. The coordinatesystem (x′, y′, z′) of the measurement is identical, with regard to thereference point, with that of the target points (x′″, y′″, z′″) as wellas with that set for the treating laser radiation. Possible deviationsfrom this identity due to instrument-specific tolerances do notinterfere strongly, because they remain constant in all cases and canthus be compensated for.

FIG. 20 shows a first regime of measurement for sensing a tissue 76.Beginning at a starting point 88, the laser focus is directed ontopoints of measurement arranged in a grid. Different types of tissue 80 aand 80 b, which contact each other at a boundary 91, lead to differentlylaser-induced signals S. Due to the evaluated signal S, each point ofmeasurement can be assigned to one type of tissue. Thus, the point ofmeasurement 82 a is located in the first type of tissue 80 a and thepoint of measurement 82 b is located in the second type of tissue 80 b.This allows an expected area 83 to be derived for the possible locationof the boundary 91. Said area is shaded in FIG. 20 from top left tobottom right. The precision with which the actual location of theboundary 91 can be indicated by the expected area 83 depends on thelateral scanning resolution 84 and on the normal scanning resolution 85.

The regime of measurement shown in FIG. 21 differs from that of FIG. 20by a higher depth resolution, i.e. by a smaller normal scanningresolution 85. The location of the boundary 91 can be determinedconsiderably more precisely for a constant lateral scanning resolution84, if the boundary 91 extends approximately parallel to the surface anddoes not have any major slopes. Therefore, depending on the assumedcourse of the boundary 91, the accuracy of measurement can be increasedby providing an asymmetrical scanning resolution, for example by anincrease in only one direction, or the number of points of measurement82 and thus the duration of measurement can be reduced by reducing thescanning resolutions in the other directions. The volume sensing has theadvantage that all structures present therein can be recognized. Thus,for example, inclusions 89 are detected equally well as the boundary 91.

However, if there is only an interest in knowing the location of theboundary 91, the regime of measurement can be changed so as to reducethe duration of measurement, as shown in FIG. 22, which illustrates athird variant of a regime of measurement, which senses the course of aboundary 91 using as few points of measurement as possible, but at highresolution. The sensing starts at a starting point 88, which is known tobe located in the first type of tissue 80 a. The focus is thendisplaced, by the predetermined scanning stepping or resolution, towardthe assumed position of the boundary 91, and the signal S is evaluated.If the focus traverses the boundary 91, the signal S changes. The depthcoordinate of the boundary 91 is thus known for this lateral position.The sensing operation is then continued at the lateral scanningresolution, starting from a new starting point 86 at the same depth.However, instead of moving the focus up or down now, alternatingmeasurements are made at increasing distances from the new startingpoint 86 above and below the expected depth position of the boundary 91.This is indicated by small bent arrows in FIG. 22.

In this manner, the depth position of the boundary 91 is determined atall lateral points of interest. The better the assumed position of theboundary 91 at the new lateral position matches the actual position, theless points of measurement will be required. Thus, general knowledgeabout the course of the boundary 91 allows a further reduction of thenumber of points of measurement required. Such general knowledge exists,for example, for the course of the Bowman's membrane, which is locatedat an approximately constant depth beneath the outer surface of thecornea and is nearly spherical. Therefore, after several points ofmeasurement, the assumed depth position of this membrane can bepredicted rather precisely, and there are only few further points ofmeasurement required in order to determine the exact position.

Also, many tissue structures may be presumed to have a smooth coursewithout depth variations at a large spatial frequency, so that a lowerlateral resolution of ca. 0.1 mm may be sufficient. An area having adiameter of 10 mm can then be detected by means of ca. 100,000 sensingpoints, at a lateral resolution of 0.1 mm and a depth resolution of 1μm.

The above-described embodiments may be employed in a particularlyadvantageous manner for the above-mentioned surgical method. To thisend, the cornea of the eye can be aspired onto a contact glass andmeasured at high depth resolution (ca. 1 μm) and low lateral resolution(e.g. 100 μm) from the epithelium to the endothelium, over the entirearea in which the surgical operation is to be effected. According to thefifth embodiment, the energy reducer 77 is switched on for this purpose,and the layers of the cornea are sensed at several positions in alateral 100 μm grid, substantially perpendicular to the surface of thecornea. In the detection beam path, multi-photon fluorescence is thendetected, for example, in a spatially resolved manner, at a suitablewavelength which is sensitive to differences in the different layersand/or boundaries. Alternatively, each of the aforementioned principlesof measurement may be applied.

A three-dimensional image of the layer of the cornea can be generatedfrom a plurality of thus obtained depth profiles. In said image, thelaterally resolved depth position of the Bowman's membrane can berecognized, which may be of importance depending on the cutting path.For the treatment, the energy reducer is removed from the beam path, sothat the desired cut is effected below the border of the epithelium whenthe scanning operation is carried out again. The epithelium thus remainslargely uninjured, so that the cut heals again after a few days.

However, it is possible not only to remove the epithelium along theBowman's membrane, but also to effect a cut located deeper down in thestroma. In doing so, the thickness of the stroma remaining on theepithelium can be precisely set by the previous measurement, thusexcluding damage to or loss of the epithelium.

1.-31. (canceled)
 32. A device for measuring an optical break-throughwhich is created in a tissue, beneath a tissue surface, by treatinglaser radiation which a laser surgical unit focuses into a treatmentfocus, said focus being located in the tissue wherein said devicecomprises a detection beam path comprising optics, wherein the opticscouple radiation emitted by the tissue from beneath the tissue surface,into the detection beam path, and a detector unit is arranged followingthe detection beam path, said detector unit generating a detectionsignal which indicates the spatial extent, position or both of theoptical break-through in the tissue.
 33. The device as claimed in claim32, further comprising an illumination radiation source, which directsillumination radiation into the tissue.
 34. The device as claimed inclaim 33, wherein the illumination radiation source supplies thetreating laser radiation.
 35. The device as claimed in claim 33, whereinthe illumination radiation source and the detection beam path are partof an interferometer structure.
 36. The device as claimed in claim 35,wherein the interferometer structure comprises a measuring arm and anadjustable reference arm and the illumination radiation has a coherencelength, in the direction of light propagation and in which theresolution at which the detection signal indicates the spatial extentdepends on the coherence length, and wherein interference appears only,if the lengths of the measuring arm and of the reference arm differ byno more than the coherence length.
 37. The device as claimed in claim35, wherein the illumination source radiation focuses the illuminationradiation into an illumination focus located in the tissue, wherein theposition of the illumination focus is adjustable to generate thedetection signal.
 38. The device as claimed in claim 37, wherein theillumination radiation is coupled into a light path of the treatinglaser radiation, and further comprising adjustable optics by which thedivergence of the illumination radiation is changeable without changingthe divergence of the treating laser radiation.
 39. The device asclaimed in claim 32, wherein the detector unit detects the radiationemitted by the tissue by confocal imaging.
 40. The device as claimed inclaim 39, wherein the detector unit generates the detection signal byadjusting the focus of the confocal imaging, preferably along a raydirection of the treating laser radiation.
 41. The device as claimed inclaim 39, wherein the optics of the detection beam path have certainlight dispersing properties, so that they comprise different focalpoints during the confocal imaging for different spectral regions,wherein the detector unit effects a spectrally selective detection ofthe radiation recorded in the confocal imaging, to generate thedetection signal.
 42. The device as claimed in claim 41, furthercomprising a multi-channel spectrometer for picking up radiation behinda pinhole.
 43. The device as claimed in claim 33, wherein the source ofillumination radiation comprises a plurality of partial radiationsources, which are individually operable and have different spectralproperties, so that spectral selective sensing is obtained bysequentially operating said partial radiation sources.
 44. The device asclaimed in claim 32, wherein the detection beam path has an optical axiswhich is located obliquely to an optical axis of the treating laserradiation or of illumination radiation.
 45. The device as claimed inclaim 33, wherein the source of illumination radiation causes a slitillumination of the tissue.
 46. The device as claimed in claim 44,further comprising a scanning unit by which the position of the opticalaxis of the detection beam path is adjustable relative to the opticalaxis of the treating laser radiation or of the illumination radiation.47. The device as claimed in claim 32, wherein the detector unitdetermines a measure of the spatial extent, the position or both ofindividual scattering centers, which are generated by the break-through.48. The device as claimed in claim 32, wherein the detection signalindicates a diameter of a plasma bubble, which was generated by anoptimal break-through.
 49. The device as claimed in claim 32, furthercomprising a scanning device for scanning the tissue.
 50. A method ofmeasuring an optical break-through which is created in a tissue, beneatha tissue surface, by treating laser radiation comprising the steps of:detecting radiation emitted by the tissue from beneath the tissuesurface; and determining a spatial extent of the optical break through,a position of the optical break through or both of the foregoing fromdetection of the emitted radiation.
 51. The method as claimed in claim50, wherein the spatial extent, the position or both of scatteringcenters generated by the optical break-through is determined.Application No.
 52. The method as claimed in claim 50, whereinobservation radiation is directed into the tissue, and radiation emittedby the tissue in the form of back-reflection is evaluated.
 53. Themethod as claimed in claim 51, wherein the radiation emitted by thetissue is interferometrically detected.
 54. The method as claimed inclaim 53, wherein a position of the radiation emitted by the tissuealong an optical axis of detection is determined from occurringinterference.
 55. The method as claimed in claim 50, wherein theradiation emitted by the tissue is detected by confocal imaging and thespatial extent is determined by adjusting a focus of said confocalimaging.
 56. The method as claimed in claim 55, wherein differentspectral focal points are generated in confocal imaging by dispersiveoptics and radiation recorded behind a pinhole is spectrally evaluated.57. The method as claimed in claim 52, wherein spectrally differentradiation is sequentially directed toward the tissue and the radiationemitted by the tissue is sequentially recorded.
 58. The method asclaimed in claim 50, wherein the emitted radiation is detected along anoptical axis which is oblique relative to an optical axis along whichthe treating laser radiation or observation radiation is directed intothe tissue.
 59. The method as claimed in claim 58, wherein the treatmentradiation is directed into the tissue as a slit-shaped beam.
 60. Themethod as claimed in claim 58, wherein an angle between the optical axisof detection and the optical axis of irradiation is adjusted to obtaininformation on the spatial extent of the interaction.
 61. The method asclaimed in claim 50, wherein a measure of the spatial extent ofindividual scattering centers of the optical break-through is generated.62. The method as claimed in claim 61, wherein a diameter of a plasmabubble is determined.